Metallic engineering materials such as stainless steels, cobalt chromium alloys (Co—Cr alloys) and nickel-titanium alloys (Ni—Ti alloys) are used in variety of medical device applications. One example of a Ni—Ti alloy is NITINOL. The use of these alloys combine mechanical, fatigue, corrosion resistance and biocompatibility properties to create devices useful in a number of medical procedures.
Since the advent of minimally invasive procedures, engineering designers have been trying to work with specific geometries whereby metallic components can be inserted through very small openings in the human body, routed to the desired location and then deployed to a useful size to fulfill the needs of the end application. One such well-known device is a coronary stent, a tubular structure used to hold open blocked or collapsed arteries. The usual method of getting stents into the space is to collapse the metal structure onto a delivery catheter having a sufficiently small overall diameter so it can be routed percutaneously to the coronary artery and expanded to a much larger diameter than the original insertion diameter.
Conventional alloys used in various medical instruments have relied on stainless steel, complex cobalt chrome alloys (such as Elgiloy™ or L605) all of which can have their mechanical properties (i.e. yield strength, ultimate tensile strength break strength, etc) modified through work hardening and annealing. These metals, even with very high yield strength, cannot sustain strains much greater than 0.2 percent (%) without suffering a permanent set. Once a bend or kink has been sustained in a medical instrument or device fabricated from one of the above alloys it is virtually impossible to straighten and remove from the body. For many permanent implants (such as stents), the device may not need to be removed and the permanent deformation may actually be useful to keep the structure in place. However, in the foregoing alloys, Hook's taw dictates that the force to deploy the implant will increase at a linear rate until the material yield point is reached and then the force will continue to increase until the material break point is reached. Additionally, these materials have significant spring back after receiving a significant deformation.
Recently medical device engineers have begun designing metallic components with shape memory alloys. In general, shape memory alloys such as NITINOL having the proper transformation temperature and processing could potentially offer two modes of shape recovery for metallic components inserted into the human body: (1) superelasticity and (2) shape memory recovery.
In the case of superelastic NITINOL the complete “elastic” recovery of strains up to 10% due to stress induced martensite (SIM) can be achieved. When superelastic NITINOL components are subjected to a stress, the strain is accommodated by austenite to martensite crystalline transformation, rather than by the mechanisms that prevail in other alloys such as slip, grain boundary sliding and dislocation motion.
Under typical process condition the stress required to form martensite will be >60,000 pounds per square inch (psi) (414 mega pascals (MPa) and the reverse transformation stress will be >30,000 psi (207 MPa). It can be observed that the reversion stress is lower than the stress at which martensite forms. These stresses are referred to as the upper and lower plateau stresses and their magnitude is dependent on the alloy composition, cold working and thermal treatment that the NITINOL, has received. As the temperature of the specimen is raised, the stress magnitude required to produce SIM is increased; however when the specimen reaches a critical temperature above (the Austenite finish temperature) Af, designated as Md, stress induced martensite cannot be formed, no matter how high the stress. In practical applications, this behavior gives rise to a limitation on using the super-elastic property since it limits the temperature range over which super-elasticity is observed; typically in the binary Ni—Ti alloys, this is a temperature range of about 60° Celsius (C) (108° Fahrenheit) (F), although a 40° C. (72° F.) range is more typical. The desirable temperature range for medical and orthodontic applications is in the region of body temperature, +10° C. to +40″C, can be achieved in these alloys.
Others have applied superelastic NITINOL to medical devices using a 50.8 atomic percent (at. %) nickel/balance titanium formulation which has been cold worked followed by a low temperature anneal to give a combination of shape memory and/or superelastic characteristics. For starting materials having an ingot Af of 0° C., this processing gives a component with an elastic range of approximately 2% to 8% over a temperature range of +15° C. to +40° C. However, the shortcomings for deployment of a stent application include: (a) the unnecessary bulk of the stent delivery system (since the delivery system must resist the high outward radial force of the compressed stent during shipping, storage and deployment); (b) the high outward radial force of the compressed stent pressing on the inside surface of the delivery sheath can add unwanted friction during deployment of the stent from the sheath; (c) at deployment the rapid stent expansion to its memorized shape can traumatize the vessel wall; and (c) the stent can cause a chronic outward force once deployed that can cause further trauma.
Jervis, U.S. Pat. No. 5,067,957, discloses that a medical device component made from superelastic NITINOL can be externally constrained outside the body via the stress induced martensite mechanism, then placed in the body and de-constrained for deployment.
Duerig, et al., U.S. Pat. No. 6,312,455, discloses a superelastic NITINOL stent for use in a lumen in a human or animal body, having a generally tubular body formed from a shape memory alloy which has been treated so that it exhibits enhanced elastic properties with a point of inflection in the stress-strain curve on loading. This enables the body to be deformed inwardly to a transversely compressed configuration for insertion into the lumen and then revert towards its initial configuration, into contact with and to support the lumen. The shape memory alloy comprises nickel, titanium and from about 3 at. % to about 20 at. %, of the alloy composition, of a ternary element selected from the group consisting of niobium, hafnium, tantalum, tungsten and gold. The ratio of the stress on loading to the stress on unloading at the respective inflection points on the loading and unloading curves is at least about 2.5:1, and the difference between the stresses on loading and unloading at the inflection points is at least about 250 MPa.
Besselink, et al., U.S. Pat. No. 6,428,634, discloses a method of processing a highly elastic stent made from a Ni—Ti—Nb based alloy which contains from about 4 to about 14 at. % Nb and in which the atomic percent ratio Ni to Ti is from about 3.8 to 1.2, comprising working the alloy sufficiently to impart a textured structure to the alloy, at a temperature below the recrystallization temperature of the alloy. Preferably, the alloy is worked at least 10%, by a technique such as rolling or drawing, or another technique which produces a similar crystal structure. The alloy has increased stiffness compared with Ni—Ti binary alloys with superelastic properties.
For the case of shape memory recovery mentioned earlier, the thermoelastic shape memory alloys can change from martensite to austenite and back again on heating and cooling over a very small temperature range, typically from 18° C. to 55° C. On cooling from the austenitic phase, often called the parent phase, martensite starts to form at a temperature designated as Ms (martensite start) and upon reaching the tower temperature, Mf (martensite finish), the alloy is completely martensitic. Upon heating from below the Mf temperature the martensite starts to revert to the austenitic structure at As, and when the temperature designated as Af is reached, the alloy is completely austenitic. These two crystalline phases have very different mechanical properties: the Young's Modulus of austenite is 12×106 psi (82,728 MPa), while that for martensite is about 4×106 psi (27576 MPa); and the yield strength, which depends on the amount of cold work the alloy is given, ranges from 28 to 100 thousand pound per square inch (ksi) (193 to 689 MPa) for austenite and from 10 to 20 ksi (68 to 138 MPa) for martensite.
Additionally, a NITINOL structure processed to exhibit shape memory and deformed in the martensitic state can recover up to 8% strain on heating to austenite. This would be an extremely handy way to deploy devices or recover accidental bending and kinking of devices in the human body if it were not for the heating and cooling extremes that must be achieved.
Simpson, et al., U.S. Pat. No. 4,770,725, discloses a Ni—Ti—Nb shape memory alloy and article, wherein niobium varies from about 2.5 to 30 at. %. Also disclosed is an article made from these nickel/titanium/niobium alloys.
Simpson, et al., U.S. Pat. No. 4,631,094, discloses a method of processing a nickel/titanium-based shape memory alloy. The method comprises over deforming the alloy so as to cause at least some amount of non-recoverable strain, temporarily expanding the transformation hysteresis by raising the austenite transformation temperature, removing the applied stress and then storing the alloy at a temperature less than the new austenite transition temperature. Simpson also discloses an article produced from this method.
Wu, et al., U.S. Pat. No. 6,053,992, discloses a mechanism that uses the shape recovery of a shape memory alloy for sealing openings or high-pressure passages. A component made of a shape memory alloy can be processed in its martensitic state to have a reduced dimension smaller than that of the opening or the passage to be sealed. Upon heating, shape recovery takes place that is associated with the reverse crystalline phase transformation of martensite. The shape recovery of the previously processed shape memory alloy component yields a diameter greater than that of the opening or passage to be sealed. The shape recovery provides the dimensional interference and force required for sealing.
Wu, et al., builds on the work of Simpson, et al., U.S. Pat. No. 4,770,725, to use both the defined chemistry and the specified process method to effect a specific heat sealing application which employs a heat activated recovery transformation.
The wide thermal hysteresis available from thermal and mechanical treatment of alloys disclosed in the literature is attractive for articles which make use of a thermally induced configuration change, since it enables an article to be stored in the deformed configuration in the martensite phase, at the same temperature at which it will then be in use, in the austenite phase. This thermal and mechanical treatment is used in a variety of industrial heat—to recover couplings and connectors (L. Mcd. Schetky, The Applications of Constrained Recovery Shape Memory Devices for Connectors, Sealing and Clamping, Proceedings Super-elastic Technologies, Pacific Grove, Calif. (1994)).
It has been reported that a reverse transformation start temperature As′ has been raised to +70° C. after specimens were deformed to 16% strain at different temperatures, where the initial states of the specimens were pure austenite phase and/or martensite phase depending upon the pre-straining temperature regime. It was found that a transformation hysteresis width of 200° C. could be attained and the reverse transformation temperatures were measured by forcing a shape-memory recovery via heating, and that up to 50% of the pre-strain could be recovered. The work was done by Xiang-Ming He, et, al, Study of the Ni41.3Ti38.7Nb20 Wide Transformation Hysteresis Shape Memory Alloy, Metallurgical and Materials Transactions, Vol. 35A, September 2004. The work cited is an optimization of various pre-straining conditions to maximize the strain recovery possible for heat recovery application.
While the wide hysteresis confers certain advantages when the thermally induced changes in configuration are to be exploited, a wide hysteresis in stress-strain behavior is generally inconsistent with the properties of an alloy that are desirable in stent or medical device applications.
Various methods have been described to deliver and implant stents. One method frequently described for delivering a stent to a desired intraluminal location includes mounting the expandable stent on an expandable member, such as a balloon, provided on the distal end of an intravascular catheter, advancing the catheter to the desired location within the patient's body lumen, inflating the balloon on the catheter to expand the stent into a permanent expanded condition and then deflating the balloon and removing the catheter. One of the difficulties encountered using other stents involved maintaining the radial rigidity needed to hold open a body lumen while at the same time maintaining the longitudinal flexibility of the stent to facilitate its delivery.
What has been needed and heretofore unavailable is a stent which has a high degree of flexibility so that it can be advanced through tortuous passageways, can be readily expanded, and yet have the mechanical strength to hold open the body lumen into which it expanded.
Thus, it is desirable to develop an alloy that is very ductile and uniquely suited for deployment of medical devices, such as stents, into the human body. In particular, for stents, it is desirable that the compressed stent maintain its shape until expanded.